Dynamic stimulus resolution adaption

ABSTRACT

Presented herein are techniques that use acoustic scene (environmental) analysis to determine the sound class of sound signals received at a hearing prosthesis and, accordingly, assess the estimated listening difficulty that the acoustic environment presents to a recipient of the hearing prosthesis. This difficulty of the recipient&#39;s listening situation can be used to adjust, adapt, or otherwise set the resolution of the electrical stimulation signals delivered to the recipient to evoke perception of the sound signals. In other words, the resolution of the electrical stimulation signals is dynamically adapted based on the present acoustic environment of the hearing prosthesis.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a continuation of U.S. application Ser. No.15/358,225, filed on Nov. 22, 2016, the contents of which are herebyincorporated in their entirety.

BACKGROUND Field of the Invention

The present invention relates generally to electrically-stimulatinghearing prostheses.

Related Art

Hearing loss, which may be due to many different causes, is generally oftwo types, conductive and/or sensorineural. Conductive hearing lossoccurs when the normal mechanical pathways of the outer and/or middleear are impeded, for example, by damage to the ossicular chain or earcanal. Sensorineural hearing loss occurs when there is damage to theinner ear, or to the nerve pathways from the inner ear to the brain.

Individuals who suffer from conductive hearing loss typically have someform of residual hearing because the hair cells in the cochlea areundamaged. As such, individuals suffering from conductive hearing losstypically receive an auditory prosthesis that generates motion of thecochlea fluid. Such auditory prostheses include, for example, acoustichearing aids, bone conduction devices, and direct acoustic stimulators.

In many people who are profoundly deaf, however, the reason for theirdeafness is sensorineural hearing loss. Those suffering from some formsof sensorineural hearing loss are unable to derive suitable benefit fromauditory prostheses that generate mechanical motion of the cochleafluid. Such individuals can benefit from implantable auditory prosthesesthat stimulate nerve cells of the recipient's auditory system in otherways (e.g., electrical, optical and the like). Cochlear implants areoften proposed when the sensorineural hearing loss is due to the absenceor destruction of the cochlea hair cells, which transduce acousticsignals into nerve impulses. An auditory brainstem stimulator is anothertype of stimulating auditory prosthesis that might also be proposed whena recipient experiences sensorineural hearing loss due to damage to theauditory nerve.

Certain individuals suffer from only partial sensorineural hearing lossand, as such, retain at least some residual hearing. These individualsmay be candidates for electro-acoustic hearing prostheses.

SUMMARY

In one aspect, a method is provided. The method comprises: receivinginput sound signals at a hearing prosthesis; determining a sound classof the input sound signals; generating, based on the input soundsignals, electrical stimulation signals for delivery to a recipient ofthe hearing prosthesis; and setting a resolution of the electricalstimulation signals based on the sound class of the input sound signals.

In another aspect, a method is provided. The method comprises: receivingsound signals at a hearing prosthesis located in an acousticenvironment; assessing, based on the sound signals, the acousticenvironment; generating electrical stimulation representative of thesound signals at a stimulus resolution that is set based on theassessment of the acoustic environment; and delivering the electricalstimulation signals to the recipient.

In another aspect, a hearing prosthesis is provided. The hearingprosthesis comprises: one or more sound input elements configured toreceive sound signals; a sound processing path configured to convert thesound signals into one or more output signals for use in deliveringelectrical stimulation to a recipient; and a stimulus resolutionadaption module configured to set a resolution of the electricalstimulation based on a sound class of the sound signals.

BRIEF DESCRIPTION OF THE DRAWINGS

Embodiments of the present invention are described herein in conjunctionwith the accompanying drawings, in which:

FIG. 1A is a schematic diagram illustrating a cochlear implant system inaccordance with embodiments presented herein;

FIG. 1B is a block diagram of a totally implantable cochlear implantsystem in accordance with embodiments presented herein;

FIG. 2 is a graph illustrating various phases of an idealized actionpotential as the potential passes through a nerve cell;

FIG. 3 is a functional block diagram illustrating a hearing prosthesisaccordance with embodiments presented herein;

FIG. 4 is a flow diagram in accordance with embodiments presentedherein;

FIGS. 5A, 5B, 5C, 5D, and 5E are schematic diagrams illustrating thespatial resolution of electrical stimulation signals in accordance withembodiments presented herein;

FIG. 6 is a block diagram of a sound processing unit in accordance withembodiments presented herein;

FIG. 7 is a flowchart of a method in accordance with embodimentspresented herein; and

FIG. 8 is a flowchart of another method in accordance with embodimentspresented herein.

DETAILED DESCRIPTION

Embodiments of the present invention are generally directed to the useof acoustic scene (environmental) analysis to determine the sound classof sound signals received at a hearing prosthesis and, accordingly,assess the estimated listening difficulty that the acoustic environmentpresents to a recipient of the hearing prosthesis. This difficulty ofthe recipient's listening situation can be used to adjust, adapt, orotherwise set the resolution (e.g., through multipolar widening and/orfocusing) of the electrical stimulation signals delivered to therecipient to evoke perception of the sound signals. In other words, theresolution of the electrical stimulation signals is dynamically adaptedbased on the present acoustic environment of the hearing prosthesis.This dynamic adaption of the stimulation resolution may optimize thetradeoff between power consumption and hearing performance.

There are a number of different types of hearing prostheses in whichembodiments of the present invention may be implemented. However, merelyfor ease of illustration, the techniques presented herein are primarilydescribed with reference to one type of hearing prosthesis, namely acochlear implant. However, it is to be appreciated that the techniquespresented herein may be used in other hearing prostheses, such asauditory brainstem stimulators, electro-acoustic hearing prostheses,bimodal hearing prostheses, etc.

FIG. 1A is a schematic diagram of an exemplary cochlear implant 100configured to implement embodiments of the present invention. Thecochlear implant 100 comprises an external component 102 and aninternal/implantable component 104.

The external component 102 is directly or indirectly attached to thebody of the recipient and typically comprises an external coil 106 and,generally, a magnet (not shown in FIG. 1A) fixed relative to theexternal coil 106. The external component 102 also comprises one or moresound input elements 108 (e.g., microphones, telecoils, etc.) fordetecting/receiving input sound signals, and a sound processing unit112. The sound processing unit 112 includes, for example, one or morebatteries (not shown in FIG. 1A) and a sound processor (also not shownin FIG. 1A). The sound processor is configured to process electricalsignals generated by a sound input element 108 that is positioned, inthe depicted embodiment, by auricle 110 of the recipient. The soundprocessor provides the processed signals to external coil 106 via, forexample, a cable (not shown in FIG. 1A).

The implantable component 104 comprises an implant body 114, a leadregion 116, and an elongate intra-cochlear stimulating assembly 118. Theimplant body 114 comprises a stimulator unit 120, aninternal/implantable coil 122, and an internal receiver/transceiver unit124, sometimes referred to herein as transceiver unit 124. Thetransceiver unit 124 is connected to the implantable coil 122 and,generally, a magnet (not shown) fixed relative to the internal coil 122.

The magnets in the external component 102 and implantable component 104facilitate the operational alignment of the external coil 106 with theimplantable coil 122. The operational alignment of the coils enables theimplantable coil 122 to transmit/receive power and data to/from theexternal coil 106. More specifically, in certain examples, external coil106 transmits electrical signals (e.g., power and stimulation data) toimplantable coil 122 via a radio frequency (RF) link. Implantable coil122 is typically a wire antenna coil comprised of multiple turns ofelectrically insulated single-strand or multi-strand platinum or goldwire. The electrical insulation of implantable coil 122 is provided by aflexible molding (e.g., silicone molding). In use, transceiver unit 124may be positioned in a recess of the temporal bone of the recipient.Various other types of energy transfer, such as infrared (IR),electromagnetic, capacitive and inductive transfer, may be used totransfer the power and/or data from an external device to a cochlearimplant and, as such, FIG. 1A illustrates only one example arrangement.

Elongate stimulating assembly 118 is configured to be at least partiallyimplanted in cochlea 130 and includes a plurality of longitudinallyspaced intra-cochlear electrical stimulating contacts (electrodes) 128that collectively form a contact array 126. Stimulating assembly 118extends through an opening in the cochlea 130 (e.g., cochleostomy 132,the round window 134, etc.) and has a proximal end connected tostimulator unit 120 via lead region 116 that extends through mastoidbone 119. Lead region 116 couples the stimulating assembly 118 toimplant body 114 and, more particularly, stimulator unit 120.

In general, the sound processor in sound processing unit 112 isconfigured to execute sound processing and coding to convert a detectedsound into a coded signal that represents the detected sound signals.These encoded data are sometimes referred to herein as processed soundsignals and are sent to the implantable component 104. The stimulatorunit 120 is configured to utilize the processed sound signals togenerate electrical stimulation signals that are delivered to therecipient's cochlea via one or more stimulation channels. In this way,cochlear implant stimulates the recipient's auditory nerve cells,bypassing absent or defective hair cells that normally transduceacoustic vibrations into neural activity.

FIG. 1A illustrates an arrangement in which the cochlear implant 100includes an external component. However, it is to be appreciated thatembodiments of the present invention may be implemented in cochlearimplant systems having alternative arrangements. For example, FIG. 1B isa functional block diagram of an exemplary totally implantable cochlearimplant 200 configured to implement embodiments of the presentinvention. Since the cochlear implant 200 is totally implantable, allcomponents of cochlear implant 200 are configured to be implanted underskin/tissue 205 of a recipient. Because all components are implantable,cochlear implant 200 operates, for at least a finite period of time,without the need of an external device. An external device 202 can beused to, for example, charge the internal power source (battery) 207.External device 202 may be a dedicated charger or a conventionalcochlear implant sound processor.

Cochlear implant 200 includes an implant body (main implantablecomponent) 214 and an implantable microphone 208, an elongateintra-cochlear stimulating assembly 118 as described above withreference to FIG. 1A. The microphone 208 may be disposed in, orelectrically connected to, the implant body 214. The implant body 214further comprises an internal transceiver unit 223, a sound processor227, a stimulator unit 120 as described with reference to FIG. 1A, andthe battery 207.

The sound processor 227 is configured to execute sound processing andcoding to convert received/detected sound signals (e.g., received bymicrophone 208) into processed sound signals.

The transceiver unit 223 permits cochlear implant 200 to receive and/ortransmit signals to external device 202. For example, transceiver unit223 may be configured to transcutaneously receive power and/or data fromexternal device 202. However, as used herein, transceiver unit 223refers to any collection of one or more implanted components which formpart of a transcutaneous energy transfer system. Further, transceiverunit 223 includes any number of component(s) which receive and/ortransmit data or power, such as, for example a coil for a magneticinductive arrangement, an antenna for an alternative RF system,capacitive plates, or any other suitable arrangement.

As noted above, FIG. 1A illustrates an embodiment in which the externalcomponent 102 includes the sound processor. As such, in the illustrativearrangement of FIG. 1A, processed sound signals are provided to theimplanted stimulator unit 120 via the RF link between the external coil106 and the internal coil 122. However, in the embodiment of FIG. 1B,the sound processor 227 is implanted in the recipient. As such, in theembodiments of FIG. 1B, the processed sound signals do not traverse theRF link, but instead are provided directly to the stimulator unit 120.

The human auditory system is composed of many structural components,some of which are connected extensively by bundles of nerve cells(neurons). Each nerve cell has a cell membrane which acts as a barrierto prevent intercellular fluid from mixing with extracellular fluid. Theintercellular and extracellular fluids have different concentrations ofions, which leads to a difference in charge between the fluids. Thisdifference in charge across the cell membrane is referred to herein asthe membrane potential (Vm) of the nerve cell. Nerve cells use membranepotentials to transmit signals between different parts of the auditorysystem.

In nerve cells that are at rest (i.e., not transmitting a nerve signal)the membrane potential is referred to as the resting potential of thenerve cell. Upon receipt of a stimulus, the electrical properties of anerve cell membrane are subjected to abrupt changes, referred to hereinas a nerve action potential, or simply action potential. The actionpotential represents the transient depolarization and repolarization ofthe nerve cell membrane. The action potential causes electrical signaltransmission along the conductive core (axon) of a nerve cell. Signalsmay be then transmitted along a group of nerve cells via suchpropagating action potentials.

FIG. 2 illustrates various phases of an idealized action potential 142as the potential passes through a nerve cell. The action potential ispresented as membrane voltage in millivolts (mV) versus time. Themembrane voltages and times shown in FIG. 2 are for illustrationpurposes only and the actual values may vary depending on theindividual. Prior to application of a stimulus 144 to the nerve cell,the resting potential of the nerve cell is approximately −70 mV.Stimulus 144 is applied at a first time. In normal hearing, thisstimulus is provided by movement of the hair cells of the cochlea.Movement of these hair cells results in the release of neurotransmitterinto the synaptic cleft, which in return leads to action potentials inindividual auditory nerve fibers. In cochlear implants, the stimulus 144is an electrical stimulation signal (electrical stimulation).

Following application of stimulus 144, the nerve cell begins todepolarize. Depolarization of the nerve cell refers to the fact that thevoltage of the cell becomes more positive following stimulus 144. Whenthe membrane of the nerve cell becomes depolarized beyond the cell'scritical threshold, the nerve cell undergoes an action potential. Thisaction potential is sometimes referred to as the “firing” or“activation” of the nerve cell. As used herein, the critical thresholdof a nerve cell, group of nerve cells, etc. refers to the thresholdlevel at which the nerve cell, group of nerve cells, etc. will undergoan action potential. In the example illustrated in FIG. 2, the criticalthreshold level for firing of the nerve cell is approximately −50 mV.The critical threshold and other transitions may be different forvarious recipients and so the values provided in FIG. 2 are merelyillustrative.

The course of the illustrative action potential in the nerve cell can begenerally divided into five phases. These five phases are shown in FIG.2 as a rising phase 145, a peak phase 146, a falling phase 147, anundershoot phase 148, and finally a refractory phase (period) 149.During rising phase 145, the membrane voltage continues to depolarizeand the point at which depolarization ceases is shown as peak phase 146.In the example of FIG. 2, at this peak phase 146, the membrane voltagereaches a maximum value of approximately 40 mV.

Following peak phase 146, the action potential undergoes falling phase147. During falling phase 147, the membrane voltage becomes increasinglymore negative, sometimes referred to as hyperpolarization of the nervecell. This hyperpolarization causes the membrane voltage to temporarilybecome more negatively charged than when the nerve cell is at rest. Thisphase is referred to as the undershoot phase 148 of action potential142. Following this undershoot phase 148, there is a time period duringwhich it is impossible or difficult for the nerve cells to fire. Thistime period is referred to as the refractory phase (period) 149.

As noted above, the nerve cell must obtain a membrane voltage above acritical threshold before the nerve cell may fire/activate. The numberof nerve cells that fire in response to electrical stimulation (current)can affect the “resolution” of the electrical stimulation. As usedherein, the resolution of the electrical stimulation or the “stimulusresolution” refers to the amount of acoustic detail (i.e., the spectraland/or temporal detail from the input acoustic sound signal(s)) that isdelivered by the electrical stimulation at the implanted electrodes inthe cochlea and, in turn, received by the primary auditory neurons(spiral ganglion cells). As described further below, electricalstimulation has a number of characteristics/attributes that control thestimulus resolution. These attributes include for example, the spatialattributes of the electrical stimulation, temporal attributes of theelectrical stimulation, frequency attributes of the electricalstimulation, instantaneous spectral bandwidth attributes of theelectrical stimulation, etc. The spatial attributes of the electricalstimulation control the width along the frequency axis (i.e., along thebasilar membrane) of an area of activated nerve cells in response todelivered stimulation, sometimes referred to herein as the “spatialresolution” of the electrical stimulation. The temporal attributes referto the temporal coding of the electrical stimulation, such as the pulserate, sometimes referred to herein as the “temporal resolution” of theelectrical stimulation. The frequency attributes refer to the frequencyanalysis of the acoustic input by the filter bank, for example thenumber and sharpness of the filters in the filter bank, sometimesreferred herein as the “frequency resolution” of the electricalstimulation. The instantaneous spectral bandwidth attributes refer tothe proportion of the analyzed spectrum that is delivered via electricalstimulation, such as the number of channels stimulated out of the totalnumber of channels in each stimulation frame.

The spatial resolution of electrical stimulation may be controlled, forexample, through the use of different electrode configurations for agiven stimulation channel to activate nerve cell regions of differentwidths. Monopolar stimulation, for instance, is an electrodeconfiguration where for a given stimulation channel the current is“sourced” via one of the intra-cochlea electrodes 128, but the currentis “sunk” by an electrode outside of the cochlea, sometimes referred toas the extra-cochlear electrode (ECE) 139 (FIGS. 1A and 1B). Monopolarstimulation typically exhibits a large degree of current spread (i.e.,wide stimulation pattern) and, accordingly, has a low spatialresolution. Other types of electrode configurations, such as bipolar,tripolar, focused multi-polar (FMP), a.k.a. “phased-array” stimulation,etc. typically reduce the size of an excited neural population by“sourcing” the current via one or more of the intra-cochlear electrodes128, while also “sinking” the current via one or more other proximateintra-cochlear electrodes. Bipolar, tripolar, focused multi-polar andother types of electrode configurations that both source and sinkcurrent via intra-cochlear electrodes are generally and collectivelyreferred to herein as “focused” stimulation. Focused stimulationtypically exhibits a smaller degree of current spread (i.e., narrowstimulation pattern) when compared to monopolar stimulation and,accordingly, has a higher spatial resolution than monopolar stimulation.Likewise, other types of electrode configurations, such as doubleelectrode mode, virtual channels, wide channels, defocused multi-polar,etc. typically increase the size of an excited neural population by“sourcing” the current via multiple neighboring intra-cochlearelectrodes.

The cochlea is tonotopically mapped, that is, partitioned into regionseach responsive to sound signals in a particular frequency range. Ingeneral, the basal region of the cochlea is responsive to higherfrequency sounds, while the more apical regions of the cochlea areresponsive to lower frequencies. The tonopotic nature of the cochlea isleveraged in cochlear implants such that specific acoustic frequenciesare allocated to the electrodes 128 of the stimulating assembly 118 thatare positioned close to the corresponding tonotopic region of thecochlea (i.e., the region of the cochlea that would naturally bestimulated in acoustic hearing by the acoustic frequency). That is, in acochlear implant, specific frequency bands are each mapped to a set ofone or more electrodes that are used to stimulate a selected (target)population of cochlea nerve cells. The frequency bands and associatedelectrodes form a stimulation channel that delivers stimulation signalsto the recipient.

In general, it is desirable for a stimulation channel to stimulate onlya narrow region of neurons such that the resulting neural responses fromneighboring stimulation channels have minimal overlap. Accordingly, theideal stimulation strategy in a cochlear implant would use focusedstimulation channels to evoke perception of all sound signals at anygiven time. Such a strategy would, ideally, enable each stimulationchannel to stimulate a discrete tonotopic region of the cochlea tobetter mimic natural hearing and enable better perception of the detailsof the sound signals. The present inventor has realized that, althoughfocused stimulation generally improves hearing performance, thisimproved hearing performance comes at the cost of significant increasedpower consumption, added delays to the processing path, and increasedcomplexity, etc. relative to the use of only monopolar stimulation.Additionally, the present inventor has realized that not all listeningsituations benefit from the increased fidelity offered by focusedstimulation as different listening situations present varying levels ofdifficulty to cochlear implant recipients. For example, understandingspeech in a quiet room is easier than understanding the same speech in abusy restaurant with many competing speakers. Accordingly, the presentinventor has realized that recipients benefit more or less from thedetails of sound presented using increased stimulus resolution indifferent environments.

In accordance with the embodiments presented herein, a hearingprosthesis is configured to analyze received sound signals to determinethe primary or main sound “class” of the sound signals. In general, thesound class provides an indication of the difficulty/complexity of arecipient's listening situation/environment (i.e., the environment inwhich the prosthesis is currently/presently located). Based on the soundclass of the sound signals, the hearing prosthesis is configured to setthe stimulus resolution of the electrical stimulation signals that aredelivered to the recipient to evoke perception of the sound signals. Thestimulus resolution is set in a manner that optimizes the tradeoffbetween hearing performance (e.g., increased fidelity) and powerconsumption (e.g., battery life). The hearing prosthesis uses higherresolution stimulation (i.e., stimulation that provides relatively moreacoustic detail) in more challenging listening situations with increasedexpected listening effort, and uses lower resolution stimulation (i.e.,stimulation that provides relatively less acoustic detail) in easierlistening situations with lower expected listening effort. Since thereis limited power available in a cochlear implant, it is thereforeadvantageous to adapt the stimulation resolution depending on thelistening situation in order to optimize the stimulus resolution for thebest overall hearing performance within the long-term power budget.

In accordance with the embodiments presented herein, the spatialresolution (i.e., the spatial attributes of the electrical stimulation)can be increased, for example, through use of focused stimulation.Conversely, the spatial resolution can be lowered, for example, throughthe use of monopolar, or wide/defocused, stimulation. These decreases inthe stimulation resolution have the benefit of lower power consumptionand lower complexity, but they also sacrifice listening fidelity (e.g.,loss of sound details). In addition, the temporal resolution (i.e., thetemporal attributes of the electrical stimulation) can be varied, forexample, by changing the rate of the current pulses forming theelectrical stimulation. Higher pulse rates offer higher temporalresolution and use more power, while lower pulse rates offer lowertemporal resolution and are more power efficient.

Consequently, the stimulus resolution can be varied with differingassociated power costs and, in certain situations, the techniquespresented herein purposely downgrade hearing performance (e.g., speechperception) to reduce power consumption. However, this downgrade inhearing performance is dynamically activated only in listeningsituations where the recipient likely does not have difficultyunderstanding/perceiving the sound signals with lower stimulusresolution (e.g., monopolar stimulation, defocused stimulation, etc.)and/or does not need the details provided by high resolution (e.g.,focused stimulation).

FIG. 3 is a schematic diagram illustrating the general signal processingpath 150 of a cochlear implant, such as cochlear implant 100, inaccordance with embodiments presented herein. As noted, the cochlearimplant 100 comprises one or more sound input elements 108. In theexample of FIG. 3, the sound input elements 108 comprise two microphones109 and at least one auxiliary input 111 (e.g., an audio input port, acable port, a telecoil, a wireless transceiver, etc.). If not already inan electrical form, sound input elements 108 convert received/inputsound signals into electrical signals 153, referred to herein aselectrical input signals, that represent the received sound signals. Asshown in FIG. 3, the electrical input signals 153 are provided to apre-filterbank processing module 154.

The pre-filterbank processing module 154 is configured to, as needed,combine the electrical input signals 153 received from the sound inputelements 108 and prepare those signals for subsequent processing. Thepre-filterbank processing module 154 then generates a pre-filteredoutput signal 155 that, as described further below, is the basis offurther processing operations. The pre-filtered output signal 155represents the collective sound signals received at the sound inputelements 108 at a given point in time.

The cochlear implant 100 is generally configured to execute soundprocessing and coding to convert the pre-filtered output signal 155 intooutput signals that represent electrical stimulation for delivery to therecipient. As such, the sound processing path 150 comprises a filterbankmodule (filterbank) 156, a post-filterbank processing module 158, achannel selection module 160, and a channel mapping and encoding module162.

In operation, the pre-filtered output signal 155 generated by thepre-filterbank processing module 154 is provided to the filterbankmodule 156. The filterbank module 156 generates a suitable set ofbandwidth limited channels, or frequency bins, that each includes aspectral component of the received sound signals. That is, thefilterbank module 156 comprises a plurality of band-pass filters thatseparate the pre-filtered output signal 155 into multiplecomponents/channels, each one carrying a single frequency sub-band ofthe original signal (i.e., frequency components of the received soundssignal).

The channels created by the filterbank module 156 are sometimes referredto herein as sound processing channels, and the sound signal componentswithin each of the sound processing channels are sometimes referred toherein as band-pass filtered signals or channelized signals. Theband-pass filtered or channelized signals created by the filterbankmodule 156 are processed (e.g., modified/adjusted) as they pass throughthe sound processing path 150. As such, the band-pass filtered orchannelized signals are referred to differently at different stages ofthe sound processing path 150. However, it will be appreciated thatreference herein to a band-pass filtered signal or a channelized signalmay refer to the spectral component of the received sound signals at anypoint within the sound processing path 150 (e.g., pre-processed,processed, selected, etc.).

At the output of the filterbank module 156, the channelized signals areinitially referred to herein as pre-processed signals 157. The number‘m’ of channels and pre-processed signals 157 generated by thefilterbank module 156 may depend on a number of different factorsincluding, but not limited to, implant design, number of activeelectrodes, coding strategy, and/or recipient preference(s). In certainarrangements, twenty-two (22) channelized signals are created and thesound processing path 150 is said to include 22 channels.

The pre-processed signals 157 are provided to the post-filterbankprocessing module 158. The post-filterbank processing module 158 isconfigured to perform a number of sound processing operations on thepre-processed signals 157. These sound processing operations include,for example, channelized gain adjustments for hearing loss compensation(e.g., gain adjustments to one or more discrete frequency ranges of thesound signals), noise reduction operations, speech enhancementoperations, etc., in one or more of the channels. After performing thesound processing operations, the post-filterbank processing module 158outputs a plurality of processed channelized signals 159.

In the specific arrangement of FIG. 3, the sound processing path 150includes a channel selection module 160. The channel selection module160 is configured to perform a channel selection process to select,according to one or more selection rules, which of the ‘m’ channelsshould be use in hearing compensation. The signals selected at channelselection module 160 are represented in FIG. 3 by arrow 161 and arereferred to herein as selected channelized signals or, more simply,selected signals.

In the embodiment of FIG. 3, the channel selection module 156 selects asubset ‘n’ of the ‘m’ processed channelized signals 159 for use ingeneration of electrical stimulation for delivery to a recipient (i.e.,the sound processing channels are reduced from ‘m’ channels to ‘n’channels). In one specific example, the ‘n’ largest amplitude channels(maxima) from the ‘m’ available combined channel signals/masker signalsis made, with ‘m’ and ‘n’ being programmable during initial fitting,and/or operation of the prosthesis. It is to be appreciated thatdifferent channel selection methods could be used, and are not limitedto maxima selection.

It is also to be appreciated that, in certain embodiments, the channelselection module 160 may be omitted. For example, certain arrangementsmay use a continuous interleaved sampling (CIS), CIS-based, or othernon-channel selection sound coding strategy.

The sound processing path 150 also comprises the channel mapping module162. The channel mapping module 162 is configured to map the amplitudesof the selected signals 161 (or the processed channelized signals 159 inembodiments that do not include channel selection) into a set of outputsignals (e.g., stimulation commands) that represent the attributes ofthe electrical stimulation signals that are to be delivered to therecipient so as to evoke perception of at least a portion of thereceived sound signals. This channel mapping may include, for example,threshold and comfort level mapping, dynamic range adjustments (e.g.,compression), volume adjustments, etc., and may encompass selection ofvarious sequential and/or simultaneous stimulation strategies.

In the embodiment of FIG. 3, the set of stimulation commands thatrepresent the electrical stimulation signals are encoded fortranscutaneous transmission (e.g., via an RF link) to an implantablecomponent 104 (FIGS. 1A and 1B). This encoding is performed, in thespecific example of FIG. 3, at the channel mapping module 162. As such,channel mapping module 162 is sometimes referred to herein as a channelmapping and encoding module and operates as an output block configuredto convert the plurality of channelized signals into a plurality ofoutput signals 163.

Also shown in FIG. 3 are a sound classification module 164, a batterymonitoring module 166, and a stimulus resolution adaption module 167.The sound classification module 164 is configured to evaluate/analyzethe input sound signals and determine the sound class of the soundsignals. That is, the sound classification module 164 is configured touse the received sound signals to “classify” the ambient soundenvironment and/or the sound signals into one or more sound categories(i.e., determine the input signal type). The sound classes/categoriesmay include, but are not limited to, “Speech,” “Noise,” “Speech+Noise,”“Music,” and “Quiet.” As described further below, the soundclassification module 164 may also estimate the signal-to-noise ratio(SNR) of the sound signals. In one example, the operations of the soundclassification module 164 are performed using the pre-filtered outputsignal 155 generated by the pre-filterbank processing module 154.

The sound classification module 164 generates sound classificationinformation/data 165 that is provided to the stimulus resolutionadaptation module 168. The sound classification data 165 represents thesound class of the sound signals and, in certain examples, the SNR ofthe sound signals. Based on the sound classification data 165, thestimulus resolution adaptation module 168 is configured to determine alevel of stimulus resolution that should be used in deliveringelectrical stimulation signals to represent (evoke perception of) thesound signals. The level of stimulus resolution that should be used indelivering electrical stimulation signals is sometimes referred toherein as the “target” stimulus resolution.

The stimulus resolution adaptation module 168 is configured to adjustone or more operations performed in the sound processing path 150 so asto achieve the target stimulus resolution (i.e., adapt the resolution ofthe electrical stimulation that is delivered to the recipient). Thestimulus resolution adaptation module 168 may adjust operations of thefilterbank module 156, the post-filterbank processing module 158, thechannel selection module 160, and/or the mapping and encoding module 162to generate output signals representative of electrical stimulationsignals having the target stimulus resolution.

The stimulus resolution adaptation module 168 may adjust operations ofthe sound processing path 150 at a number of different time scales. Forexample, the stimulus resolution adaptation module 168 may determine thetarget stimulus resolution and make corresponding processing adjusts inresponse to a triggering event, such as the detection of a change in thelistening environment (e.g., when the sound classification data 165indicates the cochlear implant 100 is in a listening environment that isdifferent from the previous listening environment). Alternatively, thestimulus resolution adaptation module 168 can determine the targetstimulus resolution and make corresponding processing adjustssubstantially continuously, periodically (e.g., every 1 second, every 5seconds, etc.,), etc.

FIG. 3 illustrates an arrangement in which the cochlear implant 100 alsocomprises a battery monitoring module 166. The battery monitoring module166 is configured to monitor the charge status of the battery/batteries(e.g., monitor charge level, remaining battery life, etc.) and providebattery information 167 to the stimulus resolution adaptation module168. In addition to the sound classification data 165, the stimulusresolution adaptation module 168 may also use the battery information167 to determine the target stimulus resolution and make correspondingprocessing adjusts to the sound processing path operations. For example,if the battery information 167 indicates that the cochlear implantbattery/batteries are below a threshold charge level (e.g., below 20%charge), the stimulus resolution adaptation module 168 can switch thesound processing path 150 to a power saving mode that uses lowerresolution (e.g., monopolar stimulation or defocused stimulation only)to conserve power.

FIG. 3 also illustrates a specific arrangement that includes one soundclassification module 164. It is to be appreciated that alternativeembodiments may make use of multiple sound classification modules. Insuch embodiments, the stimulus resolution adaption module 168 isconfigured to utilize the information from each of the multiple soundclassification modules to determine a target stimulus resolution andadapt the sound processing operations accordingly (i.e., so that theresulting stimulation has a resolution that corresponds to the targetstimulus resolution).

Although FIG. 3 illustrates a cochlear implant arrangement, it is to beappreciated that the embodiments presented herein may also beimplemented in other types of hearing prosthesis. For example, thetechniques presented herein may be used in electro-acoustic hearingprostheses that are configured to deliver both acoustical stimulationand electrical stimulation to a recipient. In such embodiments, theprosthesis would include two parallel sound processing paths, where thefirst sound processing path is an electric sound processing path(cochlear implant sound processing path) similar to that is shown inFIG. 3. In such arrangements, the second sound processing path is anacoustic sound processing path (hearing aid sound processing path) thatis configured to generate output signals for use in acousticallystimulating the recipient.

FIG. 4 is a flow diagram illustrating further details of the techniquespresented herein. For ease of illustration, FIG. 4 will be describedwith reference to the arrangement of FIG. 3.

The flow of FIG. 4 begins at 172 where the cochlear implant 100 receivesinput sound signals for analysis by the sound classification module 164.At 173, the sound classification module 164 determines the sound classof the input sound signals. FIG. 4 illustrates example sound classes 175that include: “Speech,” “Music,” “Noise,” and “Quiet.” However, it is tobe appreciated that additional sound classes are also possible.

In the example of FIG. 4, the stimulus resolution adaption module 168 isconfigured to adapt the sound processing path 150 differently dependingon the determined sound class. The different adaptions are generallyshown in FIG. 4 at blocks 177(A), 177(B), and 177(C). More specifically,FIG. 4 illustrates that for the “Music” and “Noise” sound classes, ahigher target stimulus resolution is preferred since these are moredifficult listening situations for a recipient. Accordingly, at 177(B),the stimulus resolution adaption module 168 implements one or moreadjustments to the sound processing path 150 to set a higher targetresolution for the electrical stimulation delivered to the recipient.For the “Quiet” sound class, at 176(C) the stimulus resolution adaptionmodule 168 implements one or more adjustments to the sound processingpath 150 to set a lower target stimulus resolution.

For the “Speech” sound class, a signal-to-noise ratio (SNR) of the inputsound signals is used to control the resolution. The SNR provides ameasure of how much speech compared to noise is present in the inputsound signals. Therefore, at 176, the SNR of the input sound signals isestimated and, at 176(A) the SNR is used to determine the targetstimulus resolution. In general, the higher the SNR, the lower thetarget stimulus resolution and, conversely, the lower the SNR, thehigher the target stimulus resolution. The SNR can be used to select oneof a number of discrete stimulation resolution levels, or could beapplied across a continuum. For example, in one embodiment, the stimulusresolution adaption module 168 uses only two (2) stimulus resolutionlevels for the “Speech” sound class. These two levels comprise a lowresolution setting for SNRs greater than a certain threshold (e.g.,greater than 10 dB) where speech is much stronger than the noise, and ahigh resolution setting for SNRs below the threshold (e.g., below 10dB). In other embodiments, the stimulus resolution adaption module 168may make use of a large number of different resolution levels that eachcorrelate to different SNR ranges. In these embodiments, the determinedSNR is mapped to one of the number of different ranges and, accordingly,used to select the corresponding stimulus resolution.

Also shown in FIG. 4 is the battery information 167 which may be anoptional input to the resolution adaptation blocks 177(A)-177(C). Asnoted, the battery information 167 represents the state of charge of thebattery/batteries and can be used to modify the target resolutions at177(A)-177(C). For example, if the battery charge is low, for examplebelow 20%, then a target higher resolution at 177(A) or 177(B) might belowered so that battery life is given more priority. The batteryinformation 167 could also be used to modify the SNR threshold(s) forthe speech class. For example, in a low battery situation (e.g., chargebelow a threshold level), the SNR threshold(s) for selecting between tworesolution steps could also be lowered, for example from 10 dB to 5 dB,to favor the use of lower resolution and, accordingly, conserve power.

As described elsewhere herein, the stimulus resolution adaption module168 may set or adjust various operations of the sound processing path150, such as the operations of the filterbank module 156, thepost-filterbank processing module 158, the channel selection module 160,and/or the mapping and encoding module 162, to set the stimulusresolution of the delivered electrical stimulation signals. In oneembodiment, the spatial/spectral attributes of the stimulus resolutionare set by switching between different channel/electrode configurations,such as between monopolar stimulation, wide/defocused stimulation,focused (e.g., multipolar current focusing) stimulation, etc. FIGS.5A-5E are a series of schematic diagrams illustrating exemplaryelectrode currents and stimulation patterns for five (5) differentchannel configurations. It is to be appreciated that the stimulationpatterns shown in FIGS. 5A-5C are generally illustrative and that, inpractice, the stimulation current may spread differently in differentrecipients.

Each of the FIGS. 5A-5E illustrates a plurality of electrodes shown aselectrodes 128(1)-128(9), which are spaced along the recipient's cochleafrequency axis (i.e., along the basilar membrane). FIGS. 5A-5E alsoinclude solid lines of varying lengths that extend from variouselectrodes to generally illustrate the intra-cochlear stimulationcurrent 180(A)-180(E) delivered in accordance with a particular channelconfiguration. However, it is to be appreciated that stimulation isdelivered to a recipient using charge-balanced waveforms, such asbiphasic current pulses and that the length of the solid lines extendingfrom the electrodes in each of FIGS. 5A-5E illustrates the relative“weights” that are applied to both phases of the charge-balancedwaveform at the corresponding electrode in accordance with differentchannel configurations. As described further below, the differentstimulation currents 180(A)-180(E) (i.e., different channel weightings)results in different stimulation patterns 182(A)-182(E), respectively,of voltage and neural excitation along the frequency axis of the cochlea

Referring first to FIG. 5C, shown is the use of a monopolar channelconfiguration where all of the intra-cochlear stimulation current 180(C)is delivered with the same polarity via a single electrode 128(5). Inthis embodiment, the stimulation current 180(C) is sunk by anextra-cochlear return contact which, for ease of illustration, has beenomitted from FIG. 5C. The intra-cochlear stimulation current 180(C)generates a stimulation pattern 182(C) which, as shown, spreads acrossneighboring electrodes 128(3), 128(4), 128(6), and 128(7). Thestimulation pattern 182(C) represents the spatial attributes (spatialresolution) of the monopolar channel configuration.

FIGS. 5A and 5B illustrate wide or defocused channel configurationswhere the stimulation current is split amongst an increasing number ofintracochlear electrodes and, accordingly, the width of the stimulationpatterns increases and thus provide increasingly lower spatialresolutions. In these embodiments, the stimulation current 180(A) and180(B) is again sunk by an extra-cochlear return contact which, for easeof illustration, has been omitted from FIGS. 5A and 5B.

More specifically, in FIG. 5B the stimulation current 180(B) isdelivered via three electrodes, namely electrodes 128(4), 128(5), and128(6). The intra-cochlear stimulation current 180(B) generates astimulation pattern 182(B) which, as shown, spreads across electrodes128(2)-128(8). In FIG. 5A, the stimulation current 180(A) is deliveredvia five electrodes, namely electrodes 128(3)-128(7). The intra-cochlearstimulation current 180(A) generates a stimulation pattern 182(A) which,as shown, spreads across electrodes 128(1)-128(9). In general, the widerthe stimulation pattern, the lower the spatial resolution of thestimulation signals.

FIGS. 5D and 5E illustrate focused channel configurations whereintracochlear compensation currents are added to decrease the spread ofcurrent along the frequency axis of the cochlea. The compensationcurrents are delivered with a polarity that is opposite to that of aprimary/main current. In general the more compensation current at nearbyelectrodes, the more focused the resulting stimulation pattern (i.e.,the lower the width of the stimulus patterns increase and thusincreasingly higher spatial resolutions). That is, the spatialresolution is increased by introducing increasing large compensationcurrents on electrodes surrounding the central electrode with thepositive current.

More specifically, in FIG. 5D positive stimulation current 180(D) isdelivered via electrode 128(5) and stimulation current 180(D) ofopposite polarity is delivered via the neighboring electrodes, namelyelectrodes 128(3), 128(4), 128(6), and 128(7). The intra-cochlearstimulation current 180(D) generates a stimulation pattern 182(D) which,as shown, only spreads across electrodes 128(4)-128(6). In FIG. 5E,positive stimulation current 180(E) is delivered via electrode 128(5),while stimulation current 180(E) of opposite polarity is delivered viathe neighboring electrodes, namely electrodes 128(3), 128(4), 128(6),and 128(7). The intra-cochlear stimulation current 180(E) generates astimulation pattern 182(E) which, as shown, is generally localized tothe spatial area adjacent electrode 128(5).

The difference in the stimulation patterns 182(D) and 182(E) in FIGS. 5Dand 5E, respectively, is due to the magnitudes (i.e., weighting) ofopposite polarity current delivered via the neighboring electrodes128(3), 128(4), 128(6), and 128(7). In particular, FIG. 5D illustrates apartially focused configuration where the compensation currents do notfully cancel out the main current on the central electrode and theremaining current goes to a far-field extracochlear electrode (notshown). FIG. 5E is a fully focused configuration where the compensationcurrents fully cancel out the main current on the central electrode128(5) (i.e., no far-field extracochlear electrode is needed).

As noted, FIGS. 5A-5E collectively illustrate techniques for adjustingthe spatial resolution (i.e., adjusting the spatial attributes of theelectrical stimulation) in accordance with embodiments presented herein.However, also as noted, it is to be appreciated that other methods foraltering the stimulus resolution could be used in combination with, oras an alternative to, adjustments to the spatial resolution enabled bydifferent stimulation strategies. For example, another technique foradapting the stimulus resolution includes varying the temporalresolution via pulse rate (i.e., higher pulse rates for higher temporalresolutions and lower pulse rates for lower temporal resolutions). Ingeneral, changes to the temporal resolution may be implemented in thepost-filter bank processing module 158 (e.g., during calculation of thechannel envelope signals) and/or in the mapping and encoding module 162(e.g., selection of the pulse rate).

Another technique for adapting the stimulus resolution includes varyingthe instantaneous spectral bandwidth of the stimulation by changing thenumber of maxima in the channel selection. For example, theinstantaneous bandwidth can be increased by increasing the number ofchannels selected by the channel selection module 160 and decreased bydecreasing the number of channels selected by the channel selectionmodule 160.

A still other technique for adapting the stimulus resolution includesvarying the frequency resolution. The frequency resolution of thefilterbank module 156 can be increased by, for example, in an FFTfilterbank using a higher-point FFT. The frequency resolution of thefilterbank module 156 can be decreased by, for example in an FFTfilterbank using a lower-point FFT.

Shown below is a table (Table 1) illustrating different types ofstimulus attributes and associated resolution adaptions in accordancewith embodiments presented herein.

TABLE 1 Stimulation Attribute Resolution Adaption Spatial Decreaseresolution by multipolar defocusing (split stimulation across multipleelectrodes) Spatial Increase resolution by multipolar focusing. Usecompensating currents of the opposite polarity to minimize currentspread from main current from central electrode Spatial Decreaseresolution by using parallel stimulation strategies: highest resolutionwith fully sequential stimulation of each channel, resolution is loweredby first stimulating pairs of channels and further lowered bystimulation more than two channels in parallel Temporal Decreaseresolution by lowering the pulse rate for each channel Temporal Increaseresolution by increase the pulse rate for each channel InstantaneousIncrease instantaneous spectral bandwidth by spectral bandwidthincreasing the number of channels selected during each stimulation frameInstantaneous Decrease instantaneous spectral bandwidth by spectralbandwidth decreasing the number of channels selected during eachstimulation frame Frequency Increase frequency resolution of filterbankFrequency Decrease frequency resolution of filterbank

The embodiments presented herein have been primarily described withrespect to the use of the sound class as the mechanism for determining atarget stimulus resolution. It is to be appreciated that othertechniques for determining the target stimulus resolution may be used inaccordance with embodiments presented herein. For example, a directmeasure of the degree of listening difficulty (listening effort) couldbe used to determine the target stimulus resolution. The listeningeffort may be determined, for example, using electroencephalography(EEG) (i.e., an electrophysiological monitoring method to recordelectrical activity of the brain), pupil dilation, or a physiologicalmeasure of stress (e.g., blood pressure, cortisol level, etc.). In theseembodiments, if the listening effort is determined to be relatively high(with reference to a baseline), then the resolution can be increased.Conversely, if the listening effort is determined to be relatively low(with reference to a baseline), then the resolution can be decreased.

In a further embodiment, the target stimulus resolution can bedetermined based on a direct input from the recipient. The input mayindicate the recipient's listening effort or directly indicate a desiredstimulus resolution. In certain such embodiments, the system could betrained so as to, over time, automatically adjust stimulus resolutionbased on previously received recipient inputs. For example, the hearingprosthesis may be pre-configured with certain thresholds that causechanges between different stimulus resolutions. Over time, the stimulusresolution adaption module can adjust these thresholds using therecipient inputs

FIG. 6 is a schematic block diagram illustrating an arrangement for asound processing unit, such as sound processing unit 112, in accordancewith an embodiment of the present invention. As shown, the soundprocessing unit 112 includes one or more processors 184 and a memory185. The memory 185 includes sound processor logic 186, soundclassification logic 188, battery monitoring logic 189, and stimulusresolution adaption logic 190.

The memory 185 may be read only memory (ROM), random access memory(RAM), or another type of physical/tangible memory storage device. Thus,in general, the memory 185 may comprise one or more tangible(non-transitory) computer readable storage media (e.g., a memory device)encoded with software comprising computer executable instructions andwhen the software is executed (by the one or more processors 184) it isoperable to perform the operations described herein with reference tothe sound processor, sound classification module 164, battery monitoringmodule 166, and stimulus resolution adaption module 168.

FIG. 6 illustrates software implementations for the sound processor, thesound classification module 164, and the stimulus resolution adaptionmodule 168. However, it is to be appreciated that one or more operationsassociated with the sound processor, the sound classification module164, the battery monitoring module 166, and the stimulus resolutionadaption module 168 may be partially or fully implemented with digitallogic gates in one or more application-specific integrated circuits(ASICs).

Merely for ease of illustration, the sound classification module 164 andthe stimulus resolution adaption module 168 have been shown anddescribed as elements that are separate from the sound processor. It isto be appreciated that the functionality of the sound classificationmodule 164 and the stimulus resolution adaption module 168 may beincorporated into the sound processor.

FIG. 7 is a flowchart illustrating a method 700 in accordance withembodiments presented herein. Method 700 begins at 702 where a hearingprosthesis receives input sound signals. At 704, the hearing prosthesisdetermines a sound class of the input sound signals and, at 706, thehearing prosthesis generates, for delivery to a recipient of the hearingprosthesis, electrical stimulation signals that are representative ofthe input sound signals. At 708, a resolution of the electricalstimulation signals is set based on the sound class of the input soundsignals.

FIG. 8 is a flowchart illustrating a method 800 in accordance withembodiments presented herein. Method 800 begins at 802 where a hearingprosthesis located in an acoustic environment receives sound signals. At804, the hearing prosthesis assesses the acoustic environment based onthe sound signals. At 806, the hearing prosthesis generates electricalstimulation representative of the sound signals at a stimulus resolutionthat is set based on the assessment of the acoustic environment. At 808,the electrical stimulation signals are delivered to a recipient of thehearing prosthesis.

As described in detail above, presented herein are techniques thatanalyze the acoustic scene/environment of a hearing prosthesis and,accordingly, adjust, adapt, or otherwise set the resolution ofelectrical stimulation based on the acoustic environment (e.g., based onan estimated listening difficulty that the acoustic environment presentsto a recipient of the hearing prosthesis). The techniques presentedherein leverage the idea that there are many listening situations thatare not difficult for recipients and, in such situations, power shouldnot be wasted to transmit the most accurate neural representationpossible. Likewise, in more challenging listening situations that aremore taxing for recipients, it may be beneficial to use more power inorder to create a more accurate neural activation pattern that lessensthe listening burden on the recipient. Accordingly, the techniquespresented optimize power consumption and hearing performance based onthe listening situation.

It is to be appreciated that the above described embodiments are notmutually exclusive and that the various embodiments can be combined invarious manners and arrangements.

The invention described and claimed herein is not to be limited in scopeby the specific preferred embodiments herein disclosed, since theseembodiments are intended as illustrations, and not limitations, ofseveral aspects of the invention. Any equivalent embodiments areintended to be within the scope of this invention. Indeed, variousmodifications of the invention in addition to those shown and describedherein will become apparent to those skilled in the art from theforegoing description. Such modifications are also intended to fallwithin the scope of the appended claims.

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 21. A method, comprising: receiving input sound signals at ahearing prosthesis, wherein the hearing prosthesis comprises at leastone battery; determining a sound class environment of the input soundsignals; determining a charge level of the at least one battery;generating, based on the input sound signals, electrical stimulationsignals for delivery to a recipient of the hearing prosthesis via theone or more stimulation channels; and selecting, based on the soundenvironment of the sound signals and the charge level of the at leastone battery, an electrode configuration for use in delivering theelectrical stimulation signals to the recipient via the one or morestimulation channels.
 22. The method of claim 21, wherein the electrodeconfiguration controls, for a given stimulation channel, a size of anarea of nerve cells activated in response to stimulation signalsdelivered via the given stimulation channel.
 23. The method of claim 21,wherein the electrode configuration is a first electrode configurationselected from among a plurality of electrode configurations eachconfigured to activate nerve cell areas of different widths.
 24. Themethod of claim 21, wherein selecting an electrode configuration for usein delivering the electrical stimulation signals to the recipient viathe one or more stimulation channels comprises: selecting a focusedelectrode channel configuration for delivery of the electricalstimulation signals to the recipient.
 25. The method of claim 21,wherein selecting an electrode configuration for use in delivering theelectrical stimulation signals to the recipient via the one or morestimulation channels comprises: selecting a defocused electrode channelconfiguration for delivery of the electrical stimulation signals to therecipient.
 26. The method of claim 21, further comprising: setting atemporal resolution of the electrical stimulation signals based on thesound environment of the input sound signal and the charge level of theat least one battery.
 27. The method of claim 21, further comprising:setting a frequency resolution of the electrical stimulation signalsbased on the sound environment of the input sound signal the chargelevel of the at least one battery.
 28. The method of claim 21, furthercomprising: setting an instantaneous spectral bandwidth of theelectrical stimulation signals by changing the number of selectedchannels in a channel selection process based on the sound environmentof the input sound signal the charge level of the at least one battery.29. The method of claim 21, wherein determining a sound environment ofthe input sound signals comprises: determining the presence of speech inthe sound signals; and classifying the sound signals as speech signals.30. The method of claim 29, further comprising: determining asignal-to-noise ratio of the speech signals; and selecting the electrodeconfiguration further based on the signal-to-noise ratio of the speechsignals.
 31. A hearing prosthesis, comprising: one or more sound inputelements configured to receive sound signals; at least one battery; asound classification module configured to determine a sound environmentof the received sound signals; a battery monitoring module configured todetermine a charge level of the at least one battery; a sound processingpath configured to convert the sound signals into one or more outputsignals for use in delivering electrical stimulation to a recipient viaone or more stimulation channels; and a stimulus resolution adaptionmodule configured to select, based on the sound environment of the soundsignals and the charge level of the at least one battery, an electrodeconfiguration for use in delivering the electrical stimulation to therecipient via the one or more stimulation channels.
 32. The hearingprosthesis of claim 31, wherein the electrode configuration controls,for a given stimulation channel, a size of an area of nerve cellsactivated in response to stimulation signals delivered via the givenstimulation channel.
 33. The hearing prosthesis of claim 31, wherein theelectrode configuration controls, for a given stimulation channel, adegree of current spread associated with stimulation signals deliveredvia the given stimulation channel.
 34. The hearing prosthesis of claim31, wherein the electrode configuration is a first electrodeconfiguration selected from among a plurality of electrodeconfigurations each configured to activate nerve cell regions ofdifferent widths.
 35. The method of claim 31, wherein the soundclassification module is configured to: determine a presence of speechin the sound signals; classify the sound signals as speech signals; anddetermine a signal-to-noise ratio of the speech signals.
 36. The hearingprosthesis of claim 35, wherein the stimulus resolution adaption moduleis further configured to: select the electrode configuration for use indelivering the electrical stimulation to the recipient via the one ormore stimulation channels further based on the signal-to-noise ratio ofthe speech signals.
 37. The hearing prosthesis of claim 31, wherein thehearing prosthesis is a cochlear implant.
 38. The hearing prosthesis ofclaim 31, wherein the hearing prosthesis is an electro-acoustic hearingprosthesis.
 39. The hearing prosthesis of claim 31, wherein to selectthe electrode configuration for use in delivering the electricalstimulation to the recipient via the one or more stimulation channels,the stimulus resolution adaption module is configured to: select, basedon the sound environment of the sound signals and the charge level ofthe at least one battery, a focused electrode channel configuration fordelivery of the electrical stimulation to the recipient via the one ormore stimulation channels.
 40. The hearing prosthesis of claim 31,wherein to select the electrode configuration for use in delivering theelectrical stimulation to the recipient via the one or more stimulationchannels, the stimulus resolution adaption module is configured to:select from a group of electrode configurations consisting of: monpolarstimulation, bipolar stimulation, multipolar stimulation, and focusedmultipolar stimulation.